The present embodiments relate to a method for the control of a magnetic resonance system.
Magnetic resonance tomography is a method that has become widespread for obtaining images of the interior of a body. With this method, the body that is to be examined is subjected to a relatively high basic magnetic field. A high-frequency excitation signal (e.g., the B1 field) is transmitted by a transmitter antenna arrangement, as a result of which the nuclear spins of specific atoms that are resonantly excited by the high-frequency field are tipped through a specific flip angle in relation to the magnetic field lines of the basic magnetic field. The high-frequency signal (e.g., the “magnetic resonance signal”), which is radiated at the relaxation of the nuclear spins, is received by suitable receiver antenna arrangements. The raw data acquired in this way is used to reconstruct the image data. For the purpose of location coding, defined magnetic field gradients are superimposed in each case on the basic magnetic field during the transmitting and reading-out and receiving, respectively, of the high-frequency signals.
For the transmitting of the B1 field, a “whole body coil” (e.g., a “whole body antenna” or “body coil”) is permanently integrated in the scanner. The magnetic resonance measurement space is located in the whole body coil. The magnetic resonance measurement space may take the form of a patient tunnel. The antenna arrangements that are used to transmit the B1 field may also be used to receive the magnetic resonance signals.
In order to obtain images with a high signal-to-noise ratio (SNR), the magnetic resonance signals may be recorded with local coils or local coil arrangements including one or more local coils. The local coils or local coil arrangements are antenna systems that are located in the immediate vicinity on or beneath the patient. The signals received by the individual magnetic resonance antenna elements of the local coils, in the form of an induced voltage, are amplified by a low-noise preamplifier, and forwarded by cable to a receiver device of the magnetic resonance system. The signal-to-noise ratio is determined to a decisive degree by the precise structure of the antenna elements of the local coils and, for example, by the losses in the antenna elements. With very small antenna elements, a relatively good SNR may be attained close to the local coil. For this reason, and because of the possibility of accelerated measurement using k-space sub-scanning in parallel imaging processes (e.g., SENSE or GRAPPA), very dense receiver antenna arrangements with a large number of individual channels may be used. The individual antenna elements may have entirely different orientations relative to the transmitted B1 field. In addition to this, for example, to improve the signal-to-noise ratio, even with high-resolution images, high-field systems with 1.5 Tesla to 12 Tesla and more are used.
One problem that arises is that, even during the excitation of the nuclear spins, currents are induced by the high-frequency transmitter power in the antenna elements of the local coils. This may lead to an undesirable heating of the local coils. Accordingly, in order to reduce, when transmitting, the current in the antenna elements, which when receiving, are matched in resonance precisely to the magnetic resonance frequency and consequently also to the frequency (e.g., the Larmor frequency) of the transmitted excitation pulses, the resonance circuit of the antenna elements is opened by a likewise resonant detuning circuit. This is represented diagrammatically in an example in FIG. 1, which shows a magnetic resonance antenna element in a local coil. The antenna element 50 includes a conductor loop 51 that is interrupted at several points by shortening capacitors 52. The induced voltage is tapped using one of the shortening capacitors 52. To do this, the shortening capacitor 52 is bridged in parallel by a high-frequency switch 55 (e.g., in the form of a PIN diode 55) connected in series with an inductor 54. If the high-frequency switch 55 is opened, then the detuning circuit 53 is formed by the shortening capacitor 52 and the inductor 54. The entire magnetic resonance antenna element 50 is detuned in relation to the Larmor frequency. By way of the high-frequency switch 55, via two filter inductors 56 and a filter capacitor 57 connected parallel to the high-frequency switch 55, a tap is effected to a coil plug 58. Using the plug 58, the signal is forwarded to the receiver device of the magnetic resonance system. In the coil plug 58, for example, a pre-processing of the magnetic resonance signal is already integrated (e.g., a suitable preamplifier). If the local coil includes several such antenna elements 50, the antenna elements 50 may be compiled, for example, in the coil plug 58, in which the individual signals are then pre-amplified and corresponding individual follow-on spins in the coil plug may be conveyed away via a suitable cable. Using such a connection cable, by the imposition of a direct current, the PIN diode 55 may be switched open in order to detune the antenna element 50. The filter capacitor 57 serves to block off the high-frequency signals from the direct current.
Even with a technically optimum layout of the detuning circuit 53, the high-frequency currents incurred in the detuning circuit, when transmitting the excitation signals through the transmitter antenna system, lead to the heating of the circuit. This heat, by conduction or convection, may also lead to the heating of the surface of the local coil (e.g., the local coil housing). Specified IEC Standards allow for a maximum surface temperature of components that come in contact with the patient of 41° C. This specification is relatively demanding, since both the DC losses of the electronics (e.g., the preamplifier integrated in the coil plug, the frequency mixer, and the analog/digital converter) may contribute to the waste heat, as well as to the high-frequency-induced losses in the detuning circuit. In order to be able to meet this specification, therefore, when transmitting the high-frequency signals for the excitation of the nuclear spins, attention is to be paid not only to the maximum high-frequency load on the patient (e.g., the specific absorption rate (SAR)), but also to providing that the local coils are not overheated. In this respect, appropriate limit values are specified for the transmission of high-frequency pulses, such that these conditions are fulfilled. When local coils are being newly designed, a check is carried out to provide that the temperatures may not be reached when emitting the conventional approved high-frequency pulse sequences.
All-body antennae have been operated in a “homogeneous mode” (e.g., in a “CP mode”), in which the transmitting field is circular polarized and is present in as homogeneous manner as possible inside the patient tunnel. The field structure in this situation is largely determined by the transmitter antenna and the feed from a transmitter, in that a single high-frequency signal is produced with a defined fixed phase and amplitude relationship to all the components of the transmitter antenna. In newer systems, in order, for example, to avoid shadowing effects in the abdomen, two different polarization states may be set (e.g., a circular polarization and an elliptical polarization). The high-frequency distributions are largely homogeneous. In these cases, it is relatively easy to test the local coils and provide that the permissible maximum temperature of 41° C. is not exceeded on the surface of the local coil for the maximum possible mean transmitting fields. The decisive parameter of the transmitting field in this situation is, in addition to the spatial field structure, the “mean (B12)” (e.g., the mean square amplitude of the B1 high-frequency field over time). This value is a reference for the power that may be introduced into the local coils by the B1 transmitting field.
Magnetic resonance systems may include the transmitter antenna arrangements with a plurality of independent high-frequency transmitter channels. The individual transmitter channels may be occupied independently of one another by individual HF signals. A high-frequency pulse for the excitation includes several individual high-frequency pulses that are emitted in parallel over the different independent high-frequency transmitter channels and are then superimposed to form one common field. Such a “multi-channel pulse”, which, due to the parallel transmission of the individual pulses, is also designated as a “pTX pulse,” may be used, for example, as an excitation, refocusing, and/or inversion pulse. An antenna system with several independently controllable antenna components or transmitter channels may also be referred to as a “transmit array,” regardless of whether this involves a whole-body antenna or an antenna arrangement close to the body.
With such systems, therefore, the structure or the spatial distribution of the high-frequency field are not specified by the hardware but may be adjusted at will, or at least over broad ranges, by software. Accordingly, the systematic preliminary test of local coils for approval in such magnetic resonance systems may no longer be carried out. A test for all possible transmitting field states may be unattainable. The limit values, with which high-frequency fields are emitted, are thus to be restricted with regard to the temperatures at the local coils to such conservative limit values that, under any circumstances, the possibility of the permissible maximum temperatures being exceeded is excluded. Under certain circumstances, therefore, local coils used previously may no longer be used in newer systems, or are to undergo expensive retro-development and retro-fitting.